Ultrasound scan conversion with spatial dithering

ABSTRACT

An ultrasound imaging system includes a scan conversion process for converting ultrasound data into a standard display format conversion and can be performed on a personal computer by programming the computer to convert data from polar coordinates to cartesian coordinates suitable for display on a computer monitor. The data is provided from scan head enclosure that houses an array of ultrasonic transducers and the circuitry associated therewith, including pulse synchronizer circuitry used in the transmit mode for transmission of ultrasonic pulses and beam forming circuitry used in the receive mode to dynamically focus reflected ultrasonic signals returning from the region of interest being imaged.

CROSS REFERENCES TO RELATED APPLICATIONS

[0001] This is a Continuation application of U.S. Ser. No. 09/447,144,filed on Nov. 23, 1999 which is a Continuation application of U.S. Ser.No. 09/203,877, filed on Dec. 2, 1998 which is a Continuationapplication of International Application No. PCT/US97/24291 filed onDec. 23, 1997 which is a Continuation-in-part application of U.S. Ser.No. 08/773,647 filed on Dec. 24, 1996 which is a Continuation-in-part ofInternational Application No. PCT/US96/11166, filed on Jun. 28, 1996,which is a Continuation-in-Part application of U.S. Ser. No. 08/599,816,filed on Feb. 12, 1996, which is a Continuation-in-Part of U.S. Ser.Nos. 08/496,804 and 08/496,805 both filed on Jun. 29, 1995, the entirecontents of the above applications are being incorporated herein byreference.

BACKGROUND OF THE INVENTION

[0002] Conventional ultrasound imaging systems typically include ahand-held scan head coupled by a cable to a large rack-mounted consoleprocessing and display unit. The scan head typically includes an arrayof ultrasonic transducers which transmit ultrasonic energy into a regionbeing imaged and receive reflected ultrasonic energy returning from theregion. The transducers convert the received ultrasonic energy intolow-level electrical signals which are transferred over the cable to theprocessing unit. The processing unit applies appropriate beam formingtechniques such as dynamic focusing to combine the signals from thetransducers to generate an image of the region of interest.

[0003] Typical conventional ultrasound systems include transducer arrayshaving a plurality, for example 128, of ultrasonic transducers. Eachtransducer is associated with its own processing circuitry located inthe console processing unit. The processing circuitry typically includesdriver circuits which, in the transmit mode, send precisely timed drivepulses to the transducer to initiate transmission of the ultrasonicsignal. These transmit timing pulses are forwarded from the consoleprocessing unit along the cable to the scan head. In the receive mode,beam forming circuits of the processing circuitry introduce theappropriate delay into each low-level electrical signal from thetransducers to dynamically focus the signals such that an accurate imagecan subsequently be generated.

[0004] For phased array or curved linear scan heads, the ultrasoundsignal is received and digitized in its natural polar (r,θ) form. Fordisplay, this representation is inconvenient, so it is converted into arectangular (x,y) representation for further processing. The rectangularrepresentation is digitally corrected for the dynamic range andbrightness of various displays and hard-copy devices. The data can alsobe stored and retrieved for redisplay. In making the conversion betweenpolar and rectangular coordinates, the (x,y) values must be computedfrom the (r,θ) values because the points on the (r,θ) array and therectangular (x,y) grid are not coincident.

[0005] In prior scan conversion systems, each point on the (x,y) grid isvisited and its value is computed from the values of the two nearest 6values by linear interpolation or the four nearest neighbors on the(r,θ) array by bi-linear interpolation. This is accomplished by use of afinite state machine to generate the (x,y) traversal pattern, abi-directional shift register to hold the (r,θ) data samples in a largenumber of digital logic and memory units to control the process andensure that the correct asynchronously received samples of (r,θ) dataarrive for interpolation at the right time for each (x,y) point. Thisprior implementation can be both inflexible and unnecessarily complex.Despite the extensive control hardware, only a single path through the(x,y) array is possible.

SUMMARY OF THE INVENTION

[0006] In a preferred embodiment of the invention, scan data is directedinto a computer after beamforming and scan conversion is performed toconvert the scan data into a display format. In a preferred embodiment,scan conversion can be performed entirely using a software module on apersonal computer. Alternatively a board with additional hardware can beinserted to provide selected scan conversion functions or to perform theentire scan conversion process. For many applications, the softwaresystem is preferred as additional hardware is minimized so the personalcomputer can be a small portable platform, such as a laptop or palmtopcomputer.

[0007] Scan conversion is preferably performed using a spatial ditheringprocess described in greater detail below. Spatial dithering simplifiesthe computational requirements for scan conversion while retaining imageresolution and quality. Thus, scan conversion can be performed on apersonal computer without the need for more complex interpolationtechniques and still provide conversion at frame rates suitable for realtime ultrasound imaging.

[0008] Preferably, the scan conversion procedure includes an inputarray, a remap array, and an output array. The remap array is an arrayof indices or pointers, which is the size of the output image used todetermine where to get each pixel from the input array. The numbers ineach position in the remap array indicate where in the input data totake each pixel will go into the output array in the same position.Thus, the remap array and output array can be thought of as having thesame geometry while the input array and output array have the same typeof data, i.e., actual image data.

[0009] The input array has new data for each ultrasound frame, whichmeans that it processes the data and puts the data in the output arrayon every frame. In accordance with a preferred embodiment of theinvention, there is a new ultrasound frame approximately every 1130second. Consequently, the remap array data can be generated relativelyslowly (but still well under about one second) as long as the routineoperation of computing a new output image from a new input data set isperformed at the frame rate of approximately 30 frames per second. Thisallows a general purpose personal computer to perform the task ofgenerating the data for the remap array without compromisingperformance, but also without having to dedicate additional hardware tothe task. In a computing system having a digital signal processor (DSP),the DSP can perform the computations of the remap array.

[0010] Alternatively, certain scan conversion functions can be performedby hardware inserted into the personal computer on a circuit board. Thisboard or a card can be inserted and used as an interface to deliver datain the proper form to the PC bus controller.

BRIEF DESCRIPTION OF THE DRAWINGS

[0011] The foregoing and other objects, features and advantages of theinvention will be apparent from the following more particulardescription of preferred embodiments of the invention, as illustrated inthe accompanying drawings in which like reference characters refer tothe same parts throughout the different views. The drawings are notnecessarily to scale, emphasis instead being placed upon illustratingthe principles of the invention.

[0012]FIG. 1 is a block diagram of a conventional imaging array as usedin an ultrasound imaging system.

[0013]FIG. 2A is a schematic illustration of the relationship between alinear ultrasound transducer array and a rectangular scan region inaccordance with the present invention.

[0014]FIG. 2B is a schematic illustration of the relationship between acurved linear ultrasound transducer array and a curved scan region inaccordance with the present invention.

[0015]FIG. 2C is a schematic illustration of the relationship between alinear ultrasound transducer array and a trapezoidal scan region inaccordance with the present invention.

[0016]FIG. 2D is a schematic illustration of a phased array scan region.

[0017]FIG. 3 is a schematic pictorial view of a preferred embodiment ofthe ultrasound imaging system of the present invention.

[0018]FIG. 4A is a schematic functional block diagram of a preferredembodiment of the ultrasound imaging system of the invention.

[0019]FIG. 4B is a schematic functional block diagram of an alternativepreferred embodiment of the ultrasound imaging system of the invention.

[0020]FIG. 5A is a schematic diagram of a beamforming and filteringcircuit in accordance with the invention.

[0021]FIG. 5B is a schematic diagram of another preferred embodiment ofa beamforming and filtering circuit in accordance with the invention.

[0022]FIG. 5C is a schematic diagram of another preferred embodiment ofa beamforming and filtering circuit in accordance with the invention.

[0023]FIG. 5D is a schematic diagram of a low pass filter in accordancewith the invention.

[0024]FIG. 5E is an example of an interface circuit board in accordancewith the invention.

[0025]FIG. 5F is a preferred embodiment of an integrated beamformingcircuit in accordance with the inventions.

[0026]FIG. 6 is a graphical illustration of the passband of a filter inaccordance with the invention.

[0027]FIG. 7A is a schematic diagram of input points overlayed on adisplay.

[0028]FIG. 7B is a schematic diagram of a display of FIG. 6 having inputdata converted to pixels.

[0029]FIG. 8 is a schematic diagram of a preferred embodiment of ageneral purpose image remapping architecture.

[0030] FIGS. 9A-9B are a flow chart illustrating a remap arraycomputation technique in accordance with the invention.

[0031]FIG. 10 is a flow chart of an output frame computation engine.

[0032] FIGS. 11A-11B are schematic pictorial views of twouser-selectable display presentation formats used in the ultrasoundimaging system of the invention.

[0033] FIGS. 12A-12B are functional block diagrams of a preferredgraphical user interface.

[0034]FIG. 13 illustrates a dialog box for ultrasound image control.

[0035] FIGS. 14A-14D illustrate display boxes for entering systeminformation.

[0036] FIGS. 15A-15C illustrates additional dialog boxes for enteringprobe or FOV data.

[0037] FIGS. 15D-15J illustrate additional display and dialog boxes fora preferred embodiment of the invention.

[0038]FIG. 16 illustrates imaging and display operations of a preferredembodiment of the invention.

[0039] FIGS. 17A-17C illustrate preferred embodiments of integratedprobe systems in accordance with the invention.

[0040]FIG. 18 illustrates a 64 channel integrated controller of atransmit/receive circuit for an ultrasound system.

[0041]FIG. 19 illustrates another preferred embodiment of a transmit andreceive circuit.

[0042]FIG. 20 illustrates a Doppler Sonogram system in accordance withthe invention.

[0043]FIG. 21 illustrates a color flow map based on a fast fouriertransform pulsed Doppler processing system in accordance with theinvention.

[0044]FIG. 22 illustrates a processing system a waveform generation inaccordance with the invention.

[0045]FIG. 23 is a system for generating a color flow map in accordancewith the invention.

[0046]FIG. 24 is a process flow sequence for computing a color flow mapin accordance with the invention.

[0047]FIG. 25 is a process flow sequence for generating a color flow mapusing cross correlation method.

DETAILED DESCRIPTION OF THE INVENTION

[0048] A schematic block diagram of an imaging array 18 of Npiezoelectric ultrasonic transducers 18(1)-18(N) as used in anultrasound imaging system is shown in FIG. 1. The array of piezoelectrictransducer elements 18(1)-18(N) generate acoustic pulses which propagateinto the image target (typically a region of human tissue) ortransmitting media with a narrow beam 180. The pulses propagate as aspherical wave 185 with a roughly constant velocity. Acoustic echoes inthe form of returning signals from image points I_(p) or reflectors aredetected by the same array 18 of transducer elements, or anotherreceiving array and can be displayed in a fashion to indicate thelocation of the reflecting structure.

[0049] The acoustic echo from the image point I_(p) in the transmittingmedia reaches each transducer element 18(1)-18(N) of the receiving arrayafter various propagation times. The propagation time for eachtransducer element is different and depends on the distance between eachtransducer element and the image point I_(p). This holds true fortypical ultrasound transmitting media, i.e. soft bodily tissue, wherethe velocity of sound is at least relatively constant. Thereafter, thereceived information is displayed in a manner to indicate the locationof the reflecting structure.

[0050] In two-dimensional B-mode scanning, the pulses can be transmittedalong a number of lines-of-sight as shown in FIG. 1. If the echoes aresampled and their amplitudes are coded as brightness, a grey scale imagecan be displayed on a cathode ray tube (CRT) or monitor. An imagetypically contains 128 such scanned lines at 0.75° angular spacing,forming a 90° sector image. Because the velocity of sound in water is1.54×10⁵ cm/sec, the round-trip time to a depth of 16 cm will be 208 μs.Thus, the total time required to acquire data along 128 lines of sight(for one image) is 26.6 ms. If other signal processors in the system arefast enough to keep up with this data acquisition rate, two-dimensionalimages can be produced at rates corresponding to standard televisionvideo. For example, if the ultrasound imager is used to view reflectedor back scattered sound waves through the chest wall between a pair ofribs, the heart pumping can be imaged in real time.

[0051] The ultrasonic transmitter is typically a linear array ofpiezoelectric transducers 18(1)-18(N) (typically spaced half-wavelengthapart) for steered arrays whose elevation pattern is fixed and whoseazimuth pattern is controlled primarily by delay steering. The radiating(azimuth) beam pattern of a conventional array is controlled primarilyby applying delayed transmitting pulses to each transducer element18(1)-18(N) in such a manner that the energy from all the transmitterssummed together at the image point I_(p) produces a desired beam shape.Therefore, a time delay circuit is needed in association with eachtransducer element 18(1)-18(N) for producing the desired transmittedradiation pattern along the predetermined direction.

[0052] As previously described, the same array 18 of transducer elements18(1)-18(N) can be used for receiving the return signals. The reflectedor echoed beam energy waveform originating at the image point reacheseach transducer element after a time delay equal to the distance fromthe image point to the transducer element divided by the assumedconstant speed of the propagation of waves in the media. Similar to thetransmitting mode, this time delay is different for each transducerelement. At each receiving transducer element, these differences in pathlength should be compensated for by focusing the reflected energy ateach receiver from the particular image point for any given depth. Thedelay at each receiving element is a function of the distance measuredfrom the element to the center of the array and the viewing angulardirection measured normal to the array.

[0053] The beam forming and focusing operations involve forming a sum ofthe scattered waveforms as observed by all the transducers, but in thissum, the waveforms must be differentially delayed so they will allarrive in phase and properly weighted in the summation. Hence, a beamforming circuit is required which can apply a different delay on eachchannel, and vary that delay with time. Along a given direction, asechoes return from deeper tissue, the receiving array varies its focuscontinually with depth. This process is known as dynamic focusing.

[0054] After the received beam is formed, it is digitized in aconventional manner. The digital representation of each received pulseis a time sequence corresponding to a back-scattering cross section ofultrasonic energy returning from a field point as a function of range atthe azimuth formed by the beam. Successive pulses are pointed indifferent directions, covering a field of view from −45° to +45°. Insome systems, time averaging of data from successive observations of thesame point (referred to as persistence weighting) is used to improveimage quality.

[0055] FIGS. 2A-2D are schematic diagrams illustrating the relationshipbetween the various transducer array configurations used in the presentinvention and their corresponding scan image regions. FIG. 2A shows alinear array 18A which produces a rectangular scanning image region180A. Such an array typically includes 128 transducers.

[0056]FIG. 2B is a schematic diagram showing the relationship between acurved linear transducer array 18B and the resulting sectional curvedimage scan region 180B. Once again, the array 18B typically includes 128adjacent transducers.

[0057]FIG. 2C shows the relationship between a linear transducer array18C and a trapezoidal image region 180C. In this embodiment, the array18C is typically formed from 192 adjacent transducers, instead of 128.The linear array is used to produce the trapezoidal scan region 180C bycombining linear scanning as shown in FIG. 2A with phased arrayscanning. In one embodiment, the 64 transducers on opposite ends of thearray 18C are used in a phased array configuration to achieve the curvedangular portions of the region 180 C at its ends. The middle 64transducers are used in the linear scanning mode to complete therectangular portion of the region 180C. Thus, the trapezoidal region180C is achieved using a sub-aperture scanning approach in which only 64transducers are active at any one time. In one embodiment, adjacentgroups of 64 transducers are activated alternately. That is, first,transducers 1-64 become active. Next, transducers 64-128 become active.In the next step, transducers 2-65 are activated, and then transducers65-129 are activated. This pattern continues until transducers 128-192are activated. Next, the scanning process begins over again attransducers 1-64.

[0058]FIG. 2D shows a short linear array of transducers 18D used toperform phased array imaging in accordance with the invention. Thelinear array 18D is used via phased array beam steering processing toproduce an angular slice region 180D.

[0059]FIG. 3 is a schematic pictorial view of an ultrasound imagingsystem 10 of the present invention. The system includes a hand-held scanhead 12 coupled to a portable data processing and display unit 14 whichcan be a laptop computer. Alternatively, the data processing and displayunit 14 can include a personal computer or other computer interfaced toa CRT for providing display of ultrasound images. The data processordisplay unit 14 can also be a small, lightweight, single-piece unitsmall enough to be hand-held or worn or carried by the user. AlthoughFIG. 3 shows an external scan head, the scan head of the invention canalso be an internal scan head adapted to be inserted through a lumeninto the body for internal imaging. For example, the head can be atransesophogeal probe used for cardiac imaging.

[0060] The scan head 12 is connected to the data processor 14 by a cable16. In an alternative embodiment, the system 10 includes an interfaceunit 13 (shown in phantom) coupled between the scan head 12 and the dataprocessing and display unit 14. The interface unit 13 preferablycontains controller and processing circuitry including a digital signalprocessor (DSP). The interface unit 13 can perform required signalprocessing tasks and can provide signal outputs to the data processingunit 14 and/or scan head 12. For user with a palmtop computer, theinterface unit 13 is preferably an internal card or chip set. When usedwith a desktop or laptop computer, the interface unit 13 can instead bean external device.

[0061] The hand-held housing 12 includes a transducer section 15A and ahandle section 15B. The transducer section 15A is maintained at atemperature below 41° C. so that the portion of the housing that is incontact with the skin of the patient does not exceed this temperature.The handle section 15B does not exceed a second higher temperaturepreferably 50° C.

[0062]FIG. 4A is a schematic functional block diagram of one embodimentof the ultrasound imaging system 10 of the invention. As shown, the scanhead 12 includes an ultrasonic transducer array 18 which transmitsultrasonic signals into a region of interest or image target 11, such asa region of human tissue, and receives reflected ultrasonic signalsreturning from the image target. The scan head 12 also includestransducer driver circuitry 20 and pulse synchronization circuitry 22.The pulse synchronizer 22 forwards a series of precisely timed anddelayed pulses to high voltage driver circuits in the drivers 20. Aseach pulse is received by the drivers 20, the high-voltage drivercircuits are activated to forward a high-voltage drive signal to eachtransducer in the transducer array 18 to activate the transducer totransmit an ultrasonic signal into the image target 11.

[0063] Ultrasonic echoes reflected by the image target 11 are detectedby the ultrasonic transducers in the array 18. Each transducer convertsthe received ultrasonic signal into a representative electrical signalwhich is forwarded to preamplification circuits 24 and time-varying gaincontrol (TGC) circuitry 25. The preamp circuitry 24 sets the level ofthe electrical signals from the transducer array 18 at a level suitablefor subsequent processing, and the TGC circuitry 25 is used tocompensate for attenuation of the sound pulse as it penetrates throughhuman tissue and also drives the beam forming circuits 26 (describedbelow) to produce a line image. The conditioned electrical signals areforwarded to the beam forming circuitry 26 which introduces appropriatedifferential delay into each of the received signals to dynamicallyfocus the signals such that an accurate image can be created. Furtherdetails of the beam forming circuitry 26 and the delay circuits used tointroduce differential delay into received signals and the pulsesgenerated by the pulse synchronizer 22 are described in the incorporatedInternational Application PCT/US96/11166.

[0064] In one preferred embodiment, the dynamically focused and summedsignal is forwarded to an A/D converter 27 which digitizes the summedsignal. Digital signal data is then forwarded from the A/D 27 over thecable 16 to a color doppler processing circuit 36. It should be notedthat the A/D converter 27 is not used in an alternative embodiment inwhich the analog summed signal is sent directly over the system cable16. The digital signal is also demodulated in a demodulation circuit 28and forwarded to a scan conversion circuit 37 in the data processor anddisplay unit 14.

[0065] As also shown a scan head memory 29 stores data from a controller21 and the data processing and display unit 14. The scan head memory 29provides stored data to the pulse synchronize 22, the TGC 25 and thebeam former 26.

[0066] The scan conversion circuitry 37 converts the digitized signaldata from the beam forming circuitry 26 from polar coordinates (r,θ) torectangular coordinates (x,y). After the conversion, the rectangularcoordinate data can be forwarded to an optional post signal processingstage 30 where it is formatted for display on the display 32 or forcompression in a video compression circuit 34. The post processing 30can also be performed using the scan conversion software describedhereinafter.

[0067] Digital signal data from the A/D connector 27 is received by apulsed or continuous Doppler processor 36 in the data processor unit 14.The pulsed or continuous Doppler processor 36 generates data used toimage moving target tissue 11 such as flowing blood. In a preferredembodiment, with pulsed Doppler processing, a color flow map isgenerated. The pulsed Doppler processor 36 forwards its processed datato the scan conversion circuitry 28 where the polar coordinates of thedata are translated to rectangular coordinates suitable for display orvideo compression.

[0068] A control circuit, preferably in the form of a microprocessor 38inside of a personal computer (e.g., desktop, laptop, palmtop), controlsthe high-level operation of the ultrasound imaging system 10. Themicroprocessor 38 or a DSP initializes delay and scan conversion memory.The control circuit 38 controls the differential delays introduced inboth the pulsed synchronizer 22 and the beam forming circuitry 26 viathe scan head memory 27.

[0069] The microprocessor 38 also controls a memory 40 which stores dataused by the scan conversion circuitry 28. It will be understood that thememory 40 can be a single memory or can be multiple memory circuits. Themicroprocessor 38 also interfaces with the post signal processingcircuitry 30 and the video compression circuitry 34 to control theirindividual functions. The video compression circuitry 34 compresses datato permit transmission of the image data to remote stations for displayand analysis via a transmission channel. The transmission channel can bea modem or wireless cellular communication channel or other knowncommunication method.

[0070] The portable ultrasound imaging system 10 of the invention canpreferably be powered by a battery 44. The raw battery voltage out ofthe battery 44 drives a regulated power supply 46 which providesregulated power to all of the subsystems in the imaging system 10including those subsystems located in the scan head 12. Thus, power tothe scan head can be provided from the data processing and display unit14 over the cable 16.

[0071]FIG. 4B is a schematic functional block diagram of an alternativepreferred embodiment of the ultrasound imaging system of the invention.In a modified scan head 12′, demodulation circuitry is replaced bysoftware executed by the microprocessor 38 in a modified data processingand display unit 14′. In particular, the digital data stream from theA/D converter 27 is buffered by a FIFO memory 37. The microprocessorexecutes software instruction to demodulate, perform scan conversion,color doppler processing, post signal processing and video compression.Thus many hardware functions of FIG. 4A are replaced by software storedin memory 40 in FIG. 4B, reducing hardware size and weight requirementsfor the system 10′.

[0072] Additional preferred embodiments for beam forming circuitry ofultrasound systems are depicted in FIGS. 5A, 5B, and 5C. Each of theseimplementations requires that sampled-analog data be down-converted, ormixed, to a baseband frequency from an intermediate frequency (IF). Thedown-conversion or mixing is accomplished by first multiplying thesampled data by a complex value (represented by the complex-valuedexponential input to the multiplier stage), and then filtering the datato reject images that have been mixed to nearby frequencies. The outputsof this processing are available at a minimum output sample rate and areavailable for subsequent display or Doppler processing.

[0073] In FIG. 5A, a set of sampling circuits 56 is used to capture adata 54 represented by a packets of charge in a CCD-based processingcircuit fabricated on an integrated circuit 50. Data are placed in oneor more delay lines and output, at appropriate times using memory andcontrol circuitry 62, programmable delay circuits 58, to an optionalinterpolation filter 60. The interpolation filter can be used to providerefined estimates of the round-trip time of a sound wave and therebyprovide better focus of the returned signals from an array of sensors.In FIG. 5A, two processing channels 52, of an array of processors, aredepicted. The outputs from the interpolation filters are combined, at ananalog summing junction 66, to provide a datum of beamformed output fromthe array.

[0074] Data obtained using an ultrasound transducer resembles the outputof a modest-bandwidth signal modulated by the center frequency of thetransducer. The center frequency, or characteristic frequency, of thetransducer is equivalent to the IF. In a sample-analog system (e.g.,using CCDs), Ω=2πf_(I)/f_(s), where fi is the intermediate frequency andf_(I) is the sampling frequency. The value n corresponds to thesample-sequence number (i.e., n=0,1,2,3,4, . . . ). The outputs of themultiplier 68 are termed, in-phase (I) or quadrature (Q) samples. Ingeneral, both I and Q values will be non zero. When the IF is chosen toequal the f_(s)/4, however, the multiplier output will only produceeither I or Q values in a repeating sequence, I, Q, −I, −Q, I, Q, −I . .. In fact, the input data are only scaled by 1 and −1. Thus, if theinput data, a, are sequentially sampled at times, a[0], a[1], a[2],a[3], a[4], . . . , a[n], the output data are a[0], j*a[1], −a[2].−j*a[3], a[4], . . . , a[n], the output data are a[0], j*a[1], −a[2],−j*[3], a [4], . . .

[0075] The I and Q outputs 74, 76 are each low-pass filtered 70, 72 toreject signal images that are mixed into the baseband. The coefficientsof the low-pass filters can be designed using a least-mean square (LMSor L2-norm) or Chebyshev (L-infinity norm) criteria. In practice, it isdesirable to reduce the number of coefficients necessary to obtain adesired filter characteristic as much as possible.

[0076] An example of a CCD implementation of a low-pass filter isillustrated in FIG. 5D. The device 90 consists of a 13-state tappeddelay line with five fixed-weight multipliers 94 to implement the filtercoefficients. As can be seen in the illustration of FIG. 6, the ripplein the passband is under 0.5 dB and the stopband attenuation is lessthan −30 dB of full scale.

[0077] The output of the low-pass filters are then decimated 78 by atleast a factor of 2. Decimation greater than 2 may be warranted if thebandwidth of the ultrasound signal is bandlimited to significantly lessthan half the sampling frequency. For most ultrasound signals, adecimation factor greater than 2 is often used because the signalbandwidth is relatively narrow relative to the sampling frequency.

[0078] The order of the decimation and the low-pass filters may beinterchanged to reduce the clocking frequency of the low-pass filters.By using a filter bank, the coefficients for the I and Q low-passfilters can be chosen such that each filter only accepts every otherdatum at its input. This “alternating clock” scheme permits the layoutconstraints to be relaxed when a decimation rate of 2 is chosen. Theseconstraints can be further relaxed if the decimation factor is greaterthan 2 (i.e., when the signal bandwidth ≦≦f/2).

[0079] The down-converted output data are passed on for furtherprocessing that may include signal-envelope detection or Dopplerprocessing. For display, the signal envelope (also referred to as thesignal magnitude) is computed as the square root of the sum of thesquares of the I and Q outputs. For the case when IF=f_(s)/4, that iseither I=0 or Q=0, envelope detection becomes trivial. The I and Q dataare often the inputs to Doppler processing which also uses the signalenvelope to extract information in the positive- and/ornegative-frequency sidebands of the signal. In FIG. 5A, only onedown-conversion stage is required following the ultrasound beamforming.

[0080] In FIG. 5B, a down-conversion stage has been placed in eachprocessing channel 52 following the sampling circuits 56. Here theproduction of I and Q data 86, 88 is performed exactly as before,however, much sooner in the system. The primary advantage of thisapproach is that the data rate in each processing channel can be reducedto a minimum, based on the ultrasound signal bandwidth and hence theselection of the low-pass filter and decimation factor. In thisimplementation, all processing channels 52 will use the samecomplex-value multipliers and identical coefficients and decimationfactors in the filter stage. As in the preceding implementation,complex-valued data are delayed and interpolated to provide beamformedoutput.

[0081] The ultrasound front end depicted in FIG. 5C is nearly identicalto that in FIG. 5B. The difference is that the interpolation stage 85,87 has been removed and replaced by choosing unique values in thecomplex-valued multipliers to provide a more-precise estimate of theprocessing-channel delay. This approach has the disadvantage that theoutput of the multiplier will always exhibit I and Q values that are nonzero. This is a consequence of the varying sampling rate around the unitcircle, in a complex-plane diagram, of the multiplier input. Thus, thisapproach can provide a more precise estimate of the sample delay in eachchannel, but at the expense of producing fully complex-valued data atthe output of each processing channel. This modification may requiremore post-processing for envelope and Doppler detection than thatpresented in the previous implementations.

[0082] A preferred embodiment of a system used to interface between theoutput of the beamforning or filtering circuit and the computer is toprovide a plug in board or card (PCMCIA) for the computer.

[0083] The board 700 of FIG. 5E illustrates an embodiment in which 16bits of digital beamformed data are received over the cable from thescanhead by differential receivers 702. A clock signal is also receivedat registers 704 along with converted differential data. The first gatearray 708 converts the 16 bits to 32 bits at half the data rate. The 32bit data is clocked into the FIFO 712 which outputs add-on data 716. Thesecond gate array 710 has access to all control signals and outputs 714to the PCI bus controller. This particular example utilizes 16 bits ofdata, however, this design can also be adapted for 32 bits or more.

[0084] Alternatively, a card suitable for insertion in a slot or port ofa personal computer, laptop or palmtop computer can also be used. Inthis embodiment the differential receivers input to registers, whichdeliver data to the FIFO and then to a bus controller that is located onthe card. The output from the controller is connected directly to thePCI bus of the computer. An alternative to the use of differentialdrivers and receivers to interconnect the scan head to the interfaceboard or card is to utilize the IEEE 1394 standard cable also known as“firewire.”

[0085] An example of a preferred embodiment of an integrated beamformingcircuit 740 is illustrated in FIG. 5F. The circuit 740 includes a timingcircuit 742, and 5 delay circuits 760 attached to each side of summingcircuit 754. Each circuit 760 includes a sampling circuit 746, a CCDdelay line 752, a control and memory circuit 750, a decoder 748, and aclocking driver circuit 744. The circuitry is surrounded by contact pads756 to provide access to the chip circuitry. The integrated circuit ispreferably less than 20 square millimeters in area and can be mounted ona single board in the scan head as described in the various embodimentsset forth in the above referenced incorporated application. A sixteen,thirty two, or sixty four delay line integrated circuit can also beimplemented utilizing a similar structure.

[0086]FIG. 7A is a schematic diagram of input points overlayed on adisplay. As illustrated, input points I_(p) received from the ultrasoundbeam 180 do not exactly align with the rectangular arranged pixel pointsP of a conventional display 32. Because the display 32 can only displaypixelized data, the input points I_(p) must be converted to therectangular format.

[0087]FIG. 7B is a schematic diagram of a display of FIG. 6 having inputdata converted to pixels. As illustrated, each image point I_(p) isassigned to a respective pixel point P on the display 32 to form animage.

[0088] One purpose of scan conversion is to perform the coordinate spacetransformation required for use with scan heads that are not flatlinear, such as phased array, trapezoidal or curved linear heads. To dothis, data must be read in one order and output data must be written inanother order. Many existing systems must generate the transformationsequences on the fly, which reduces the flexibility and makestrapezoidal scan patterns more difficult.

[0089] Because scan conversion is reordering the data, it can also beused to rotate, pan and zoom the data. Rotation is useful for viewingthe image with the scan head depicted at the top, left, right, or bottomof the image, or an arbitrary angle. Zooming and panning are commonlyused to allow various parts of the image to be examined more closely.

[0090] In addition to zooming into one area of the object, it is usefulto be able to see multiple areas simultaneously in different regions ofthe screen. Often the entire image is shown on the screen but certainregions are replaced with zoomed-in-views. This feature is usuallyreferred to as “window-in-a-window.” Current high-end systems providesthis capability for one window, but it is preferred that an imagingsystem allow any number of zoomed regions, each of which having anarbitrary size and shape.

[0091] The use of irregular scan patterns can ease system design andallow greater scan head utilization. In particular, this allowsreduction or hiding of dead time associated with imaging deep zones. Inthe case of deep zone imaging, the beam is transmitted but received atsome later time after the wave has had time to travel to the maximumdepth and return. More efficient use of the system, and thus a higherframe rate or greater lateral sampling, can be obtained if other zonesare illuminated and reconstructed during this dead time. This can causethe scan pattern to become irregular (although fixed and explicitlycomputed). The flexible scan conversion described below corrects forthis automatically.

[0092]FIG. 8 is a schematic diagram of a preferred embodiment of ageneral purpose image remapping architecture. In accordance with apreferred embodiment of the invention, data is preferably broughtdirectly into the PC after beamforming and the remainder of themanipulation is performed in software. As such, additional hardware isminimized so the personal computer can be a small portable platform,such as a laptop or palmtop computer.

[0093] Preferably, there is an input array 142, a remap array 144 and anoutput array 146. The remap array 144 is an array of indices orpointers, which is the size of the output image used to determine whereto get each pixel from the input array 142. The numbers in each positionin the remap array 144 indicate where in the input data to take eachpixel which will go into the output array 146 in the same position.Thus, the remap array 144 and output array 196 can be thought of ashaving the same geometry while the input array 142 and output array 146have the same type of data, i.e., actual image data.

[0094] The input array 142 has new data for each ultrasound frame, whichmeans that it processes the data and puts the data in the output array146 on every frame. In accordance with the invention, there is a newultrasound frame at a rate of at least 20 frames per second andpreferably approximately every {fraction (1/30)} second. However, theremap array 144 is only updated when the head type or viewing parameters(i.e., zoom and pan) are updated. Consequently, the remap array 144 datacan be generated relatively slowly (but still well under about onesecond or else it can become cumbersome) as long as the routineoperation of computing a new output image from a new input data set isperformed at the frame rate of approximately 30 frames per second. Thisallows a general purpose personal computer to perform the task ofgenerating the data for the remap array 144 without compromisingperformance, but also without having to dedicate additional hardware tothe task. In- a computing system having a digital signal processor(DSP), the DSP can perform the computations of the remap array 144.

[0095] In a preferred embodiment of the invention, input memory for theinput array 142 can be either two banks of Static Random Access Memory(SRAM) or one bank of Video Random Access Memory (VRAM), where the inputis serial access and the output is random access. The VRAM bank,however, may be too slow and refresh too costly. The remap memory forthe remap array 144 is preferably sequential access memory embodied inVRAM, or Dynamic Random Access Memory (DRAM), although random accessSRAM will also work. The output memory for the output array 146 can beeither a frame buffer or a First-In First-Out (FIFO) buffer. Basically,the scan conversion is done on demand, on the fly. Scan conversion ispreferably performed by software in the PC. If scan conversion is donein hardware, however, the PC is merely storing data, thus reducingsystem complexity. Thus, an architecture in accordance with theinvention is preferably just two random access input buffers, asequential access remap buffer and small (if any) FIFO or bit ofpipelining for the output buffer. This implies the output frame bufferis in PC memory.

[0096] In accordance with a preferred embodiment of the invention, aspatial dithering technique employing error diffusion is used inultrasound scan conversion. Typical dithering is done in the pixelintensity domain. In accordance with the invention, however, ditheringis used in ultrasound scan conversion to approximate pixels in thespatial domain and not in the pixel intensity domain. Spatial ditheringis used to approximate values that fall between two input data points.This happens because only discrete radii are sampled but pixels on thedisplay screen can fall between two radii and need to be filtered.Spatial dithering must be used to interpolate between longitudinalsample points.

[0097] Recall that the remap array 144 stores the mapping of each outputpoint to an input point. The input data points are typically in polarcoordinates while the output points are in rectilinear coordinates.Although the remap array 144 merely contains indices into the inputarray 142, they can be considered to contain radius (r) and angle (θ)values. Ideally, these values have arbitrary precision and do not haveto correspond to actual sampled points. Now consider that thesearbitrary precision numbers must be converted into integer values. Theinteger radius values correspond to discrete samples that were taken andare limited by the radial sampling density of the system. The integerangle values correspond to discrete radial lines that were scanned andare thus limited by the number of scan angles. If spatial dithering isapplied, these floating point values can be mapped into fixed integervalues without having the artifacts that appear with discrete roundingwithout error diffusion.

[0098] FIGS. 9A-9B are a flow chart illustrating a remap arraycomputation technique in accordance with the invention. At step 205, thescan heads are checked to see if there has been any change. If the scanheads have been changed, processing continues to step 210 where the newhead type is configured. After step 210, or if there has been no changein the scan heads (step 205) processing continues to step 215. At step215, the display window is checked to see if there is any zooming,panning or new window-in-window feature. If so, processing continues tostep 220 where the user inputs the new viewing parameters. After step220, or if there is no window change at step 215, processing continuesto step 225 where the remap array is cleared to indicate a newrelationship between the input and output arrays.

[0099] At step 230, the program chooses a window W to process. At step235, all line error values LE and all sample error values SE areinitialized to zero. At step 240, a point counter P is initialized topoint to the top left pixel of the window W.

[0100] At step 245, the application computes a floating point linenumber L_(FP) and sample offset S_(FP) for each point in a view V. For aphased array, this would be a radius r and an angle θ. At step 250, anypreviously propagated error terms L_(E), S_(E) (discussed below) areadded to the floating point values L_(FP), S_(FP) for the point P. Atstep 255, floating point terms are rounded to the nearest integer L_(R),S_(R), which correspond to actual sampled points. At step 260, theapplication computes rounding errors as:

L _(RE) =L _(FP) −L _(R);

S _(RE) =S _(FP) −S _(R).

[0101] At step 265, the errors are propagated to the pixel points to theright, below left, below, and below right relative to the current pointP. PROPAGATE ERRORS L_(E (right)) = L_(E (right)) + L_(RE) * 7/16L_(E (below left)) = L_(E (below left)) + L_(RE) * 3/16 L_(E (below)) =L_(E (below)) + L_(RE) * 5/16 L_(E (below right)) =L_(E (below right)) + L_(RE) * 1/16 S_(E (right)) = S_(E (right)) +S_(RE) * 7/16 S_(E (below left)) = S_(E (below left)) + S_(RE) 3/16S_(E (below)) = S_(E (below)) + S_(RE) * 5/16 S_(E (below right)) =S_(E (below right)) + S_(RE) * 1/16

[0102] At step 270, the application computes a data index based on ascan data ordered index:

REMAP(P)=Index(L _(R) , S _(R)).

[0103] At step 275, a check is made to see if there are more points inthe window. If there are more points to be processed, the pointer P isincremented to the next point at step 280. Processing then returns tostep 245. Once all points in the window have been processed, processingcontinues to step 285.

[0104] At step 285, a check is made to see if there are more windows tobe processed. If so, processing returns to step 230. Otherwise,processing is done.

[0105] Because the dithering maps one source to each output pixel, thesame remapping architecture can be used to make real-time scanconversion possible in software, even on portable computers. Thus, thecomplicated dithering operation is only performed during initializationor when viewing parameters are changed. However, the benefits of thedithering are present in all the images.

[0106]FIG. 10 is a flow chart of an output frame computation engine. Atstep 305, beamforming, demodulated input data is read into memory. Atstep 310, the output pixel index P is initialized. At step 315, theoutput array is set equal to the remapped input array according to thefollowing:

OUTPUT(P)=INPUT(REMAP(P)).

[0107] At step 320, the output pixel index P is incremented. At step325, a check is done on the pixel index P to see if the image has beenformed. If not, processing returns to step 315. Once all the pixels inthe image have been computed, processing continues to step 330 where theoutput image is optionally smoothed. Finally, at step 335, the outputimage is displayed.

[0108] Although dithering does remove the mach-banding and moire patternartifacts which occur with simple rounding, dithering can introducehigh-frequency noise. It is this high-frequency noise whose averagevalue allow for the smooth transition effects. To the untrained eye,these artifacts are far less objectionable than those obtained with thesimple rounding or nearest-point case, but may be objectionable toultrasound technicians.

[0109] These artifacts can be greatly reduced or potentially eliminatedby employing a low-pass spatial filter to smooth the image after theremapping process. The filter can be a box filter or non-symmetricalfilters can be matched to a desired input resolution characteristic.Filters can be applied in the rectilinear domain that match theorientation or angle of point coordinates at the particular location.

[0110] Basically, it is desirable to have a matched filter whose extentis similar to or proportional to distances between points beingdithered. Thus, a high magnification is preferably accompanied by alarge filter with much smoothing, whereas in places with a spacing ofthe sampled radius r or angle (θ) is small (on the order of one pixel),no filtering may be required.

[0111] Because the remapping operation is basically two loads and astore, it can be performed using a standard personal computer. Theremapping algorithm when encoded in assembly language has been shown towork on a 166 MHZ Pentium-based PC to obtain very-near real-timeoperation. In addition, the demodulation has been performed on the PCwhen written in assembly language while still achieving near real-timeoperation. Text and graphics labels are preferably effected by storingfixed values or colors in the beginning of the input buffer and thenmapping to those places where those colors are to be used. If effect,shapes or text are drawn in the remap array, which will open andautomatically be overlayed on all of the images at no computationalcost.

[0112] FIGS. 11A-11B are schematic pictorial views of display formatswhich can be presented on the display 32 of the invention. Rather thandisplaying a single window of data as is done in prior ultrasoundimaging systems, the system of the present invention has multiple windowdisplay formats which can be selected by the user. FIG. 11A shows aselectable multi-window display in which three information windows arepresented simultaneously on the display. Window A shows the standardB-scan image, while window B shows an M-scan image of a Dopplertwo-dimensional color flow map. Window C is a user information windowwhich communicates command selections to the user and facilitates theuser's manual selections. FIG. 11B is a single-window optional displayin which the entire display is used to present only a B-scan image.Optionally, the display can show both the B-mode and color doppler scanssimultaneously by overlaying the two displays or by showing themside-by-side using a split screen feature.

[0113]FIG. 12 is functional block diagram of a preferred graphical userinterface. A virtual control 400 includes an ultrasound image controldisplay 410, a probe model properties display 420, and a probe specificproperties display 500. The virtual control display 400 is preferablycoded as dialog boxes in a Windows environment.

[0114]FIG. 13 illustrates a dialog box for the ultrasound image control410. Through the ultrasound image control display 410, the user canselect a probe head type 412, a zone display 414, a demodulation filter416, and an algorithm option 418. The user also can initiate theultrasound scan through this dialog box.

[0115] The probe model properties display 420 includes model type 425,safety information 430, image Integrated Pulse Amplitude (IPA) data 435,doppler IPA data 440, color IPA data 445, probe geometry 450, imagezones data 455, doppler zones data 460, color zones data 465, imageapodization 470, doppler apodization 475, and color apodization 480.These are preferably encoded as dialog boxes. Through themodel-properties dialog box 425, a user can enter general settings forthe probe model.

[0116]FIG. 14A illustrates a dialog box for entering a viewing probemodel properties. Entered parameters are downloaded to the ultrasoundprobe.

[0117]FIG. 14B illustrates a dialog box for entering and viewing safetyinformation 430. As illustrated, a user can enter general settings 432and beam width table data 434 per governing standards.

[0118]FIG. 14C illustrates a dialog box for entering and viewing imageIPA data 435. The dialog box displays beamformed output values, listedin volts as a function of image display zones for various drivevoltages. Similar dialog boxes are used to enter the doppler and colorIPA data 440, 445.

[0119]FIG. 14D illustrates a dialog box for effecting the imageapodization function 470. As illustrated, the operator can enter andview general settings 472 and vector information 474. The user canselect active elements for array windowing (or apodization).

[0120] The probe specific property display 500 includes dialog boxes forentering probe specifics 510, image Field Of-View (FOV) data 520,doppler FOV data 530, and color FOV data 540. Through the probespecifics dialog box 510, the user can enter general settings 512,imaging static information 514, doppler static information 516, and FOVsettings 518.

[0121]FIG. 15A illustrates a dialog box for entering, and viewing probespecific information. Any number of probes can be supported.

[0122]FIG. 15B-15C illustrate dialog boxes for entering image FOV data520. As illustrated, a user can enter general settings 522, breakpointPGC data 524, zone boundaries 526, and zone duration 528 data. Dialogboxes for the doppler and color FOV data displays 530,540 are similarand are all the entry of general settings 532, 542, breakpoint TGC data534, 544, and PRF data 536, 546.

[0123] FIGS. 15D-15J illustrate additional windows and control panelsfor controlling an ultrasound imaging system in accordance with theinvention. FIG. 15D shows a viewing window for the region of interestand a control panel situated side by side with the scan image. FIG. 15Eshows controls for the doppler field of view and other selectablesettings. FIG. 15F shows the color field of view controls. FIG. 15Gshows properties of the probe. FIG. 15h shows the color IPA data for aprobe. FIG. 151 shows the probe geometry settings for a linear array.FIG. 15J shows settings for doppler apodization.

[0124]FIG. 16 illustrates the zoom feature of a preferred embodiment ofthe imaging system in accordance with the invention. In this particularillustration detailed features of a phantom, or internal anatomicalfeatures 600 of a patient that are shown on screen 32, can be selectedand enlarged within or over a display window. In this particularexample, a region 602 is selected by the user and is enlarged at window604. A plurality of such regions can be simultaneously enlarged andshown on screen 32 in separate or overlying windows. If two scan headsare in use, different views can be shown at the same time, or previouslyrecorded images can be recalled from memory and displayed beside animage presented in real time.

[0125] The architecture of the integrated front-end probe approach wasdesigned to provide small size, low power consumption and maximalflexibility in scanning, including: 1) multi-zone focus on transmission;2) ability to drive a variety of probes, such as linear/curved linear,linear/trapezoidal, and sector scan; 3) ability to provide M-mode,B-mode, Color Flow Map and Doppler Sonogram displays; 4) multiple,selectable pulse shapes and frequencies; and 5) different firingsequences. Different embodiments for the integrated front-end system 700are shown in FIGS. 17A, 17B and 17C. Modules unique to this inventionare the blocks corresponding to: beamforming chip 702, transmit/receivechip 704, preamplifier/TGC chip 706.

[0126] The block labeled “front-end probe” (front-end controller)directly controls the routine operation of the ultrasound scan head bygenerating clock and control signals provided to modules 702, 704, 706and to the memory unit 708. These signals are used to assure continuousdata output and to indicate the module for which the data appearing atthe memory-unit output are intended. Higher level control of the scanhead 710, as well as initialization, data processing and displayfunctions, are provided by a general purpose host computer 720, such asa desktop PC, laptop or palmtop. Thus, the front-end controller alsointerfaces with the host computer, e.g. via PCI bus or Fire Wire 714 toallow the host to write control data into the scanhead memory unit andreceive data back. This is performed at initialization and whenever achange in parameters (such as number and/or position of zones or type ofscan head) is required when the user selects a different scanningpattern. The front-end controller also provides buffering andflow-control functions, as data from the beamformer must be sent to thehost via a bandwidth-constrained link, to prevent data loss.

[0127] The system described permits two different implementations of theColor Flow Map (CFM) and Doppler Sonogram (DS) functions. FIG. 17A showsa hardware-based 722 implementation, in which a dedicatedDoppler-processing chip is mounted on a back-end card 724 and used as aco-processor to the host computer 720 to accomplish the CFM and DScomputations. FIG. 17B shows a software implementation in which the CFMand DS computations are performed by the host computer.

[0128]FIG. 17C shows yet another system integration, in which thetransducer array and the front-end processing units are not integratedinto a single housing but are connected by coaxial cables. The front-endunits include the front-end controller, the memory and the three modules704 (transmit/receive chip), 706 (preamp/TGC chip) and 702 (thebeamforming chip) as shown in the Figure.

[0129] “FireWire” refers to IEEE standard 1394, which provideshigh-speed data transmission over a serial link. This allows use ofhigh-volume, low cost commercial parts for the interface. The standardsupports an asynchronous data transfer mode that can be used to sendcommands and configuration data to the probe head memory. It can also beused to query the status of the head and obtain additional information,such as the activation of any buttons or other input devices on thehead. Additionally, the asynchronous data transfer mode can be used todetect the type of probe head attached. An isochronous transfer mode canbe used to transfer data back from the beamformer to the host. Thesedata may come directly from the A/D or from the demodulator or somecombination. If Doppler processing is placed in the probe head, theDoppler processed data can be sent via FireWire. Alternatively the datacan be Doppler processed via software or hardware in the host. Therealso exists a wireless version of the FireWire standard, allowingcommunication via an optical link for untethered operation. This can beused to provide greater freedom when the probe head is attached to thehost using wireless FireWire.

[0130] The preamp/TGC chip as implemented consists of integrated 32parallel, low-noise, low-power, amplifier/TGC units. Each unit has 60-dBprogrammable gain, a noise voltage less than 1.5nV1{square root}{squareroot over (Hz)} and dissipates less than 11 mW per receiver channel.

[0131] As shown in FIG. 18, the multi-channel transmit/receive chipconsists of a global counter, a global memory and a bank of paralleldual-channel transmit/receiver controllers. Within each controller 740,there are local memory 745, delay comparator, frequency counter &comparator, pulse counter & Comparator, phase selector, transmit/receiveselect/demux switch (T/R switch), and level shifter units.

[0132] The global counter 742 broadcasts a master clock and bit valuesto each channel processor 740. The global memory 744 controls transmitfrequency, pulse number, pulse sequence and transmit/receive select. Thelocal delay comparator 746 provides delay selection for each channel.For example, with a 60 MHZ clock, and a 10-bit global counter, a delayof up to 17 μs can be provided for each channel. The local frequencycounter 748 provides programmable transmit frequency. A 4-bit counterwith a comparator provides up to sixteen different frequency selections.For example, using a 60-MHZ master clock, a 4-bit counter can beprogrammed to provide different transmit frequencies such as 60/2=30MHz, 60/3=20 MHz, 60/4=15 MHz, 60/5=12 MHz, 60/6=10 MHz and so on. Thelocal pulse counter 750 provides different pulse sequences. For example,a 6-bit counter with a comparator can provide programmable transmittedpulse lengths from one pulse up to 64 pulses. The locally programmablephase selector which provides sub-clock delay resolution.

[0133] While typically the period of the transmit-chip determines thedelay resolution, a technique called programmable subclock delayresolution allows the delay resolution to be more precise than the clockperiod. With programmable subclock delay resolution, the output of thefrequency counter is gated with a phase of the clock that isprogrammable on a per-channel basis. In the simplest form, a two-phaseclock is used and the output of the frequency counter is either gatedwith the asserted or deasserted clock. Alternatively, multiple skewedclocks can be used. One per channel can be selected and used to gate thecoarse timing signal from the frequency counter. For example, for a60-MHz master clock, a two-to-one phase selector provides 8-ns delayresolution and a four-to-one phase selector provides 4-ns delayresolution.

[0134] Also shown are the integrated transmit/receiver select switch754, T/R switch and the integrated high-voltage level shifter 750 forthe transmit pulses. A single-chip transmit/receive chip capable ofhandling 64 channel drivers and 32-channel receivers can be used, eachchannel having a controller as shown in FIG. 18.

[0135] In another implementation, shown in FIG. 19, the T/R select/muxswitch and the high-voltage level shifter are separated from the othercomponents 760 on a separate chip 762 to allow use of differenthigh-voltage semiconductor technologies, such as high-breakdown siliconCMOS/JFET or GaAs technology for production of these components.

[0136] The basic method for pulsed-Doppler ultrasound imaging isillustrated in FIG. 20. The waveform consists of a burst of N pulses770. After each pulse as many range (depth) samples as needed arecollected. The time evolution of the velocity distribution of materialwithin the range gate is displayed as a sonogram 772, a two-dimensionaldisplay in which the horizontal axis represents time and the verticalaxis velocity (as assessed by Doppler shift). Different regions can beinterrogated by moving the range gate and varying its size. A Dopplersonogram can be generated using single-range-gate Doppler processing, asshown in FIG. 20. The operation of this method is as follows. A sequenceof N ultrasonic pulses is transmitted at a pulse repetition frequencyf_(prf) along a given viewing angle. The return echoes are range gatedand only returns 774 from a single range bin are used, meaning that onlythe returned signals corresponding to a region at a selected distance(e.g. from depth d to d+δd) from the transducer array along the selectedviewing angle are processed to extract Doppler information. The velocityprofiles of scatterers in the selected region can be obtained bycomputing the Doppler shifts of the echoes received from the scatterers.That is, Fourier transformation 776 of the received time-domain signalprovides frequency information, including the desired Doppler, f_(d).The velocity distribution of the scatterers in the region of interestcan be obtained from the relationship: $f_{d} = {2\frac{v}{c}f_{c}}$

[0137] where c is the speed of sound in the transmitting medium andf_(c), is the center frequency of the transducer. As an example, if N=16and f_(prf)=1 KHz, the above equation can be used to generate a sonogram772 displaying 16 ms of Doppler data. If the procedure is repeated everyN/f_(prf) seconds, a continuous Doppler sonogram plot can be produced.

[0138] Another embodiment involves a pulse-Doppler process for colorflow map applications. It is clinically desirable to be able to displayflow rates and patterns over a large region in real time. One method forapproaching this task using ultrasound is called color flow mapping(CFM). Color flow mapping techniques are an extension of thesingle-gated system described above. In CFM, velocities are estimatednot only along a single direction or line segment, but over a number ofdirections (multiple scan lines) spanning a region of interest. Thevelocity information is typically color-coded (e.g. red indicates flowtoward the transducer, blue away) and superimposed over a B-mode imagethat displays the underlaying anatomy.

[0139] A color-flow map 780 based on pulsed-Doppler processing is shownin FIG. 21. The basic single-range bin system of FIG. 20 can be extendedto measure a number of range gates by sampling at different depths andretaining the samples in storage for additional processing. Note thatthis does not increase the acquisition time, as data are collected fromthe same RF line. Sweeping the beam over an area then makes it possibleto assemble an image of the velocities in a 20 region of interest. Inoperation, the data from J range bins 782 along a single direction areprocessed in parallel. After N pulse returns are processed, the outputsrepresent a J×N range-vs-Doppler distribution, which in turn can be usedto generate a J×N velocity distribution profile. The mean velocity ateach depth d_(k)k=1,2 . . . J, is used to generate a single point orcell on the color-flow map; in each cell, the standard deviation is usedto assess turbulence. If the procedure is repeated every N/f_(prf)seconds for every J range bins (e.g. spaced J/2 range bins apart) andfor every scan line in the region of interest, a 2D color-flow map plotcan be produced.

[0140] It is important to note that instead of an FFT-based computation,a cross correlation technique, as described in the publication of JorgenA. Jensen, “Estimation of Blood Velocities Using Ultrasound,” UniversityPress 1996, the contents of which is incorporated herein by reference,can also be used to produce a similar color flow map.

[0141] The range gate size and position can be determined by the user.This choice determines both the emitted pulse length and pulserepetition frequency. The size of the range gate is determined by thelength of the pulse. The pulse duration is

T _(p)=21_(g) /C=M _(fo)

[0142] if the gate length is l_(g), and M is the number of sine periods.The depth of the gate determines how quickly pulse echo lines can beacquired. The maximum rate is

f _(prf) =c/2d _(o)

[0143] where d_(o) is the distance to the gate.

[0144] The generic waveform for the pulse-Doppler ultrasound imaging isshown in FIG. 22 where the waveform consists of a burst of N pulses 800.As many as range depth samples as needed are collected following eachpulse in the burst. FIG. 22 also shows a block diagram 810 of aconventional signal processor for this imaging technique, where thereturned echoes received by each transducer are sampled and coherentlysummed prior to in-phase and quadrature demodulation. The downconverted/basebanded returns are converted to a digital representation,and then stored in a buffer memory until all the pulse returnscomprising a coherent interval are received. The N pulse returnscollected for each depth are then read from memory, a weightingsequence, v(n), is applied to control Doppler sidelobes, and an N-pointFFT is computed. During the time the depth samples from one coherentinterval are being processed through the Doppler filter, returns fromthe next coherent interval are being processed through the Dopplerfilter, returns from the next coherent interval are arriving and arestored in a second input buffer. The FFT 818 output is passed on to adisplay unit or by time averaging Doppler samples for subsequentdisplay.

[0145] The CDP device described here performs all of the functionsindicated in the dotted box of FIG. 22, except for A/D conversion, whichis not necessary because the CDP device provides the analog sampled datafunction. This CDP Pulsed-Doppler Processor (PDP) device has thecapability to compute a matrix-matrix product, and therefore has a muchbroader range of capabilities than needed to implement the functionsshown within the dotted lines.

[0146] The PDP device computes the product of two real-valued matricesby summing the outer products formed by pairing columns of the firstmatrix with corresponding rows of the second matrix.

[0147] In order to describe the application of the PDP to the Dopplerfiltering problem, we first cast the Doppler filtering equation into asum of real-valued matrix operations. The Doppler filtering isaccomplished by computing a Discrete Fourier Transform (DFT) of theweighted pulse returns for each depth of interest. If we denote thedepth-Doppler samples g(kj), where k is the Doppler index, O≦k≦N−1, andj is the depth index, then${g\left( {k,j} \right)} = {\sum\limits_{n = 0}^{n - 1}{{v(n)}{f\left( {n,j} \right)}{\exp \left( {{- {j2}}\quad \pi \quad {{kn}/N}} \right)}}}$

[0148] The weighting function can be combined with the DFT kernel toobtain a matrix of Doppler filter transform coefficients with elementsgiven by

W(k,n)=W _(k,n) =v(n)exp(−j2nkn/N)

[0149] The real and imaginary components of the Doppler filtered signalcan now be written as$g_{r,{kj}} = {\sum\limits_{n = 0}^{N = 1}\left( {{W_{r,{kn}}f_{r,{nj}}} - {W_{i,{kn}}f_{i,{nj}}}} \right)}$$g_{r,{kj}} = {\sum\limits_{n = 0}^{N = 1}\left( {{W_{r,{kn}}f_{r,{nj}}} + {W_{i,{kn}}f_{i,{nj}}}} \right)}$

[0150] In the above equations, the double-indexed variables may all beviewed as matrix indices. Therefore, in matrix representation, theDoppler filtering can be expressed as matrix product operation. It canbe seen that the PDP device can be used to perform each of the fourmatrix multiplications, thereby implementing the Doppler filteringoperation.

[0151] A block diagram of the PDP device described in this invention isshown in FIG. 22. The device includes a J-stage CCD tapped delay line, JCCD multiplying D/A converters (MDACs) J×K accumulators, a J×K Dopplersample buffer, and a parallel-in-serial out (PISO) output shiftregister. The MDACs share a common 8-bit digital input on which elementsfrom the coefficient matrix are supplied. The tapped delay line performsthe function of a sample-and hold, converting the continuous-time analoginput signal to a sampled analog signal.

[0152] A two-PDP implementation 840 for color flow mapping in aultrasound imaging system is shown in FIG. 23. In this device, duringone pulse return interval, the top PDP component computes all the termsof the form W_(k),f_(r) and W_(i)f_(r) as shown in the above, while thebottom component computes the terms of the form −W_(i)f_(i), andW_(k)f_(i). The outputs of each component are then summed to alternatelyobtain g_(r) and g_(i).

[0153] Doppler and color flow map processing involves a significantamount of computation. This processing may be accomplished in softwareusing a general-purpose microprocessor. The presence of instructionsoptimized for matrix-matrix operations, such as the Intel MMX featureset, can substantially improve performance. A software flow chart forcolor-flow map computation based on the FFT computation algorithm isshown in FIG. 24. After initialization 900, the downconverted data isobtained 902 and the pointer P is at the beginning of the scan line 904,the data is averaged and stored 906, a weighting function is applied908, the FFT is computed 910, the magnitude z(k) is computed for eachfrequency 912 followed by the computation of first and second moments914 and display thereof in color 916. The painter is incremented 918 andeach scan line is processed as needed.

[0154] A software flow chart for color-flow map computation based on thecross-correlation computation is showing in FIG. 25.

[0155] After initiation 940, the scan line data is obtained 942,followed by the range bin data 944. The cross correlation is computed946 and averaged 948, and the velocity distribution 950, first andsecond moments 952 are obtained and displayed 954. The range bin data isincreased 956 and the process repeated.

[0156] While this invention has been particularly shown and describedwith references to preferred embodiments thereof, it will be understoodby those skilled in the art that various changes in form and details maybe made therein without departing from the spirit and scope of theinvention as defined by the appended claims.

What is claimed:
 1. A method of processing image data with an ultrasoundimaging device comprising of: providing a portable ultrasound imagingsystem including a transducer array within a handheld probe, aninterface unit connected to the handheld probe with a first interface,the interface unit having a beamforming device and being connected to adata processing system with a second interface; transmitting data fromthe handheld probe to the interface unit with the first interface;performing a beamforming operation with the beamforming device in theinterface unit; and transmitting data from the interface unit to thedata processing system with the second interface such that the dataprocessing system receives a beamformed electronic representation of theregion of interest.
 2. The method of claim 1 wherein the data processingsystem further comprises a portable computer having a flat paneldisplay.
 3. The method of claim 1 further comprising generating acolored image of the object.
 4. The method of claim 1 further comprisingproviding a beamforming circuit including a programmable delay device.5. The method of claim 1 further comprising a hand-held probe having acircuit board on which circuit elements are mounted, the circuitelements including a charged coupled device integrated circuit connectedto an analog to digital converter.
 6. The method of claim 1 wherein thestep of providing a data processing system further comprises providing abattery powered portable computer having a graphical user interface. 7.The method of claim 1 further comprising displaying the image in one ofa plurality of windows on a display connected to the data processingsystem.
 8. The method of claim 1 further comprising performing a scanconversion with a remap array.
 9. The method of claim 1 wherein thefirst interface comprises an electrical cable.
 10. The method of claim 1wherein the second interface comprises an electrical cable.
 11. Themethod of claim 1 wherein the second interface comprises a wirelessconnection.
 12. A portable ultrasound system for imaging a region ofinterest comprising: a handheld probe housing in which a transducerarray is mounted; an interface unit connected to the handheld probehousing with a first interface, the interface unit including abeamforming device; and a data processing system connected to theinterface unit with a second interface such that the data processingsystem receives a beamformed representation of the region of interest.13. The system of claim 12 further comprising a flat panel displayconnected to the data processing system that displays an image of theregion of interest.
 14. The system of claim 12 wherein the probe housingfurther comprises a beamforming circuit board having a programmabledelay device.
 15. The system of claim 12 further comprising a circuitboard within the probe housing, the circuit board having a beamformingintegrated circuit mounted thereon.
 16. The system of claim 12 furthercomprising a display and a battery in the data processing system suchthat the battery provides power to the probe housing.
 17. The system ofclaim 12 further comprising a digital signal processor in the interfaceunit.
 18. The system of claim 12 wherein the data processing systemcomprises a personal computer having a graphical user interface.
 19. Thesystem of claim 12 further comprising a data transmitter that forwardsisochronous data from the interface unit to the data processing unit.20. The system of claim 12 wherein the second interface providesasynchronous data transfer between the interface unit and the dataprocessing system.
 21. The system of claim 12 wherein asynchronoussignals are transmitted from the data processing unit to the interfaceunit.
 22. The system of claim 12 wherein the transducer array comprisesa phased array device.
 23. The system of claim 12 wherein thebeamforming device comprises 32 channels or more.
 24. The system ofclaim 12 wherein the interface unit further comprises atransmit/receiver circuit and a preamplifier circuit.
 25. The system ofclaim 12 wherein the interface unit further comprises a memory and asystem controller.
 26. The system of claim 12 wherein the beamformingdevice comprises a CDP beamformer.
 27. The system of claim 12 whereinthe data processor includes a scan conversion system and a high standardhigh speed communications port.
 28. The system of claim 12 wherein thesecond interface comprises a wireless connection.
 29. The system ofclaim 12 wherein the data processing system is programmed to performscan conversion with a remap array.
 30. A portable ultrasound system forimaging a region of interest comprising: an ultrasound probe systemincluding a transducer array and a beamforming device; and a dataprocessing system connected to the ultrasound probe system with anisochronous transfer connector from the beamforming device to the dataprocessing system to provide a high-speed transmission link such thatthe data processing system receives a beamformed representation of theregion of interest.
 31. The system of claim 30 further comprising a flatpanel display connected to the data processing system that displays animage of the region of interest.
 32. The system of claim 30 wherein theprobe system further comprises a beamforming circuit having aprogrammable delay device.
 33. The system of claim 30 further comprisingan interface unit connected to a probe housing and a circuit boardwithin the interface unit, the circuit board having a beamformingintegrated circuit mounted thereon.
 34. The system of claim 33 furthercomprising a display and a battery in the data processing system suchthat the battery provides power to the probe housing.
 35. The system ofclaim 33 further comprising a digital signal processor in the interfaceunit.
 36. The system of claim 30 wherein the data processing systemcomprises a personal computer having a graphical user interface.
 37. Thesystem of claim 30 further comprising a data transmitter that forwardsisochronous data from the probe system to the data processing unit. 38.The system claim of 33 wherein the transmission link provides aconnection between the interface unit and the data processing unit. 39.The system of claim 30 wherein asynchronous signals are transmitted fromthe data processing system to the probe.
 40. The system of claim 33wherein the interface unit comprises a wireless interface.
 41. Thesystem of claim 30 wherein the transducer array comprises a phased arraydevice.
 42. The system of claim 30 wherein the probe system comprises ahandheld housing having a transducer array and an interface unit. 43.The system of claim 30 wherein the probe comprises a beamformer device.44. The system of claim 42 wherein the interface unit comprises abeamforming device.
 45. The system of claim 30 wherein the transmissionlink comprises a wireless interface.
 46. The system of claim 30 furthercomprising a first cable connecting the probe to an interface unit and asecond cable for the transmission link between the interface unit andthe data processor.